1. Field of the Invention
The present invention is directed to a method for generating a signal corresponding to the respiration volume per unit of time of a patient, of the type wherein a measured impedance signal is acquired with an electrode arrangement arranged in the region of the heart and a signal component correlated to the respiration of the patient is filtered out of the measured impedance signal and a selected characteristic of the filtered signal is analyzed to obtain a measurement of the respiration volume per unit of time.
2. Description Of the Prior Art
A method of this type is disclosed in U.S. Pat. No. 4,901,725, wherein a measured impedance signal of the heart of a patient is derived with an electrode arrangement of a heart pacemaker and with an impedance measuring means in the heart pacemaker, this measured impedance signal changing dependent on the pumping activity of the heart as well as dependent on intrathoracal pressure fluctuations externally acting on the heart and produced by the respiration and by movements of the patient. A signal component correlated to the respiration is filtered out of the measured impedance signal by band-pass filtering, with frequency parts of the measured impedance signal below 0.05 Hz and above 0.8 Hz are suppressed. Zero-axis crossings of the filtered-out signal component are detected with a zero-axis crossing detector, and a quantity dependent on the amount of the signal component is acquired at every zero-axis crossing, this quantity being utilized for the continuous formation of an average corresponding to the respiration volume per unit of time during a prescribed time interval. To this end, the signal component obtained by filtering is sampled and supplied to a zero-axis crossing detector as well as to an amplitude averaging unit. The amplitude averaging unit averages the amplitude of the samples over a duration corresponding to a few breaths to form an amplitude average of the signal component. It is assumed that the amplitude average of the signal component corresponds to the average volume per breath. At every detected zero-axis crossing during the course of the signal component, the momentary amplitude average of the signal component is supplied to a further averaging unit. This generates a further average value, based on the amplitude average of the signal component again identified at every zero-axis crossing, and that corresponds to the average volume per breath, according to the frequency of the zero-axis crossings that identify the respiratory rate. The further average thus corresponds to the product of volume per breath and respiratory rate, and thus to the respiration volume per time unit. The respiration volume per time unit identified in this way is utilized for the frequency control of the heart pacemaker.
As already mentioned, it is assumed in the known method that the amplitude average of the signal component that is filtered out of the measured impedance signal and is correlated to respiration corresponds to the average volume per breath. Such an assumption, however, is only true when the curve of the signal component is at least approximately sinusoidal and lies symmetrically relative to the zero line. In practice, however, the signal components, despite the filtering, are superimposed with signal parts that are based on the heart activity and on movements of the patient. These signal parts have in fact been reduced in signal height as a consequence of the filtering but can produce zero-axis crossings in the course of the signal component that are erroneously detected as inspiration or expiration of the measured impedance signal and are evaluated with the calculated value for the average volume per breath in the known method. A further source of error is that the lower limit frequency for the band-pass filtering of the measured impedance signal is fixed at an extremely low value, 0.05 Hz in the known method, in view of the lowest respiration rate to be anticipated, so that the zero line of the signal component can fluctuate considerably during the time required for the formation of the amplitude average of the signal component that corresponds to only a few breaths, and therefore the "real" zero line cannot be acquired. The formation of the amplitude of the signal component and their averaging therefore proceeds based on an artificially-set zero line, resulting in deviations of the artificially-set zero line from the real zero line lead to a constant which is not representative of respiration entering into the averaging, which is expressed as an error in the identification of the volume per breath.